1. Field of the Invention
The present invention relates to a coil for detecting a magnetic resonance signal induced in a sample and, more particularly, to a coil that is immune to environmental noise and also, is less sensitive to the effects of nearby electrically conducting media.
2. Description of the Related Art
Magnetic resonance is useful to detect the presence of a specific substance in a sample. For example, generally, radio frequency (RF) radiation at a particular frequency will induce a magnetic resonance signal in a specific substance, but not in other substances. Therefore, the induced magnetic resonance signal can be detected to thereby indicate the presence of the specific substance.
It is common to detect a magnetic resonance signal by placing a sample to be measured in a tuned, electronically resonant tank circuit. Then, the response of the tank circuit to the electromotive force produced by nuclear or electronic spins in the sample is measured. With Nuclear Magnetic Resonance (NMR) or Nuclear Quadrupole Resonance (NQR), the sample is placed in or near an inductor, commonly referred to as a coil, that detects AC magnetic fields. The inductance of the coil is tuned with a parallel and/or series capacitance to make the circuit electrically resonant at the measurement frequency. One or more additional reactive impedances (inductors or capacitors) are typically added to adjust the resistive impedance at resonance to a particular value which optimizes the detection sensitivity.
FIG. 1 is a diagram illustrating an example of a conventional magnetic resonance apparatus. Referring now to FIG. 1, a transmitter 20 and a receiver 22 are connected to a probe 24 through a transmit/receive (T/R) switch 26. Probe 24 includes a coil 28, forming part of a resonant, tuned tank circuit with various other inductors L and capacitors C. To detect the presence of a target substance, T/R switch 26 connects transmitter 20 to probe 24 while disconnecting receiver 22 from probe 24. Then, transmitter 20 generates a pulse and supplies the pulse to probe 24. As an example, in NQR, the pulse is formed from an RF signal having a frequency corresponding to the resonance signal of the target substance which is intended to be detected. Probe 24 receives the pulse, which causes coil 28 to store (RF) energy.
If a sample (not illustrated) is appropriately placed near or inside coil 28, the stored RF energy will cause a corresponding RF magnetic field to irradiate the sample. If the sample includes the target substance, the RF magnetic field will induce a magnetic resonance signal in the target substance. For example, if the apparatus operates under the principles of NMR, then an appropriate NMR resonance signal will be induced. If the apparatus operates under the principles of NQR, then an appropriate NQR resonance signal will be induced.
After the sample is irradiated with the RF magnetic field, T/R switch 26 connects receiver 22 to probe 24 while disconnecting transmitter 20 from probe 24. Coil 28 then detects the resonance induced in the target substance, and probe 24 produces a corresponding output signal. The output signal of probe 24 is received and analyzed by receiver 22, to confirm the presence and/or measure the quantity of the target substance in the sample.
FIG. 1 is only one example of a magnetic resonance apparatus. For example, FIG. 1 illustrates T/R switch 26 to connect transmitter 20 and receiver 22 to the same probe 24. However, instead, a transmitter and receiver can each have a separate, dedicated probe together with a switch or gate for protecting the receiver while the transmitter is ON.
FIG. 2 is a diagram illustrating a simple, conventional coil which can be used in a probe. Referring now to FIG. 2, a coil 29 typically forms a loop. Typically, a tuning capacitance C and a matching capacitance C' are also provided.
In magnetic resonance the signal-to-noise ratio (SNR) is determined, in part, by the noise contributions and the quality factor (Q) of the receiver coil. It is well-known that random thermal noise contributions typically arise from Johnson noise in the RF inspection coil and the first amplifier in the receiver. A further noise contribution is from extraneous environmental noise.
It is also well-known in magnetic resonance that the Q of a receiver coil is determined not only by resistive loss in the windings of coil itself but also by loss in nearby electrically conducting samples that can dissipate energy from the receiver coil. As the SNR typically varies as Q.sup.1/2, such electrical loss in the sample leads to a reduction in SNR. For example, in MRI, the main source of electrical loss can come from the patient, and not the receiver coil windings, as the water in the body has an electrical conductivity comparable to sea water. In NQR landmine detection it is found that wet soils also present significant electrical loading to the receiver coil, leading to a decreased coil Q and decreased SNR.
In many applications in MRI and also landmine detection, a surface coil is used for the inspection. The larger the surface coil, the more the receiver Q is decreased. Indeed, in MRI, a system designer conventionally chooses as the receiver coil the smallest surface coil that will "cover" the region of interest.
For a conventional simple circular coil of radius R on the surface of a large conducting volume, we find that approximately 65% of the electrical loss from this volume arises from regions that are deeper than a distance R below the surface. However, because the RF magnetic field falls off so rapidly with distance, this coil is very inefficient at receiving NMR or NQR signals much beyond a distance R. Hence, most of the loss comes from a region beyond the actual region that can be imaged (in MRI) or inspected (in NQR landmine detection).
Accordingly, it is desired to reduce the environmental noise pickup and reduce the electrical loss due to the proximity of a conducting medium. Approaches for achieving this may employ detector (or transmitter) coils that create electromagnetic fields designed to (i) reject electromagnetic environmental noise that does not vary across the dimensions of the coil and (ii) to couple strongly to signals that arise near the surface of a sample, and more weakly with those deeper inside a sample.
In most NQR, NMR and EPR applications, a common coil would typically be used as both a receiver coil and a transmitter coil, though this is not essential. In MRI separate coils would typically be used.
Even though the RF coil used as a detector in magnetic resonance is not at all optimized for the detection of radio signals, nonetheless the detection coil can act as a (rather inefficient) radio receiving antenna. This radio interference, either from radio stations, or other RF noise in the relevant frequency range can easily overwhelm the magnetic resonance signals of interest. Therefore, conventionally, an external RF shield is used to surround the coil to prevent such radio signals or other electrical and magnetic environmental noise from being picked up by the coil. Unfortunately, an external RF shield is impractical in many applications, and can also increase the cost and size of a system.
There are two main strategies to remove external interference: shielding that prevents the offending interference from entering the detection coil and `balancing` in which the interference enters but is arranged to cancel itself based on some property of the detection coil and the external interference. A gradiometer is one example of the latter approach.
In principle, a gradiometer can be designed to respond to any order of the spatial derivative of the electric or magnetic field. Here we concentrate on a conventional (linear or first order) gradiometer, that is (primarily) sensitive to the first derivative of the field, and correspondingly insensitive to the component that does not vary. Hence such a conventional gradiometer provides a means to cancel fields that have wavelengths much larger than the characteristic size of the gradiometer, thereby reducing noise pickup from distant sources. (For example, the free space wavelength of a 1 MHz signal is 300 meters.). Generally, a conventional (linear) magnetic field gradiometer is formed of two loops which are spatially removed from each other and have currents flowing in opposite directions.
FIG. 3 is a diagram illustrating a conventional magnetic field gradiometer. Referring now to FIG. 3, a conventional magnetic field gradiometer has a conductor 30 which forms two loops 32 and 34 wound in the opposite sense. That is, the direction of the current changes in coils 32 and 34, so that magnetic fields generated by coils 32 and 34 are opposite each other.
While a conventional magnetic field gradiometer may reduce environmental magnetic field noise pick-up, the amount of reduction is often not enough for many applications. Therefore, an external RF shield must often be used with a conventional magnetic field gradiometer to reduce environmental electrical noise pick-up.
Instead of using an external RF shield, various conventional coil designs reduce environmental noise pick-up by providing "internal" shielding. Here, internal shielding refers to the coil being electrically balanced to cancel out electrical noise. It has long been known that such electrical balancing can reduce the interaction between the coil and electric fields from its surroundings, so that there is no net electric dipole formed between the coil and the surroundings.
A Stensgaard split-shield resonator coil is a type of coil which provides internal shielding to reduce environmental noise pick-up from electric fields.
FIG. 4 is a diagram illustrating a Stensgaard split-shield resonator coil. Referring now to FIG. 4, a coaxial cable 36 has a center conductor 38 and a shield 40, and forms a loop. Shield 40 is split at a position 42 half way around the loop. An RF feed is provided via center conductor 38. The resulting currents which flow on the outside of shield 40 are automatically electrically balanced due to geometry. Connecting portion 43 indicates that the shield 40 of each end of transmission line 36 is connected together near the feed point. A tuning capacitance C, a matching capacitance C' and a terminating impedance Z are typically used. This type of split-shield resonator is considered to be an optimized, internally shielded (that is, self-shielded) loop resonator, and is well suited for NMR medical imaging applications.
From measurements performed by the inventors of the present invention with a 20 cm diameter version of a Stensgaard split-shield resonator coil, the noise level without external RF shielding was approximately 20 to 25 dB larger than the noise level with external RF shielding. Further, the noise without RF shielding could not be significantly reduced with additional shielding from external electric fields. This result indicates that while the Stensgaard split-shield resonator removes electrical noise pick-up, it is still prone to magnetic interference.